In the computed tomography (CT) system, X-rays, which are radiated from the focal spot of an X-ray resource like a point, pass through a subject to be checked. After penetrating the subject, the X-rays are attenuated in two-dimensional distribution and measured by a detector, and then its distributed data are processed by a computer to generate a corresponding tomography image.
In the current medical computed tomography system, a solid detector is mainly used for X-ray measurement. Such a solid detector is often an array of an X-ray photoelectric receivers, for example, 16×16 or 16×32 pixel matrix, formed by arranging a plurality of photoelectric semiconductor units in a grid matrix.
In the ideal condition, an X-ray travels in a straight line. The data that shall be collected on each pixel of an X-ray detector correspond to the attenuation of the X-rays which penetrate the subject to be checked in the straight path from the focal spot to this pixel. The X-ray that is radiated from the focal spot to the X-ray detector in a straight line is referred to as a primary ray.
In the practical use, when an X-ray emanated from the focal spot passes through the subject to be checked to a surface of the detector, it is inevitable that the X-ray interacts with the subject so that the primary ray is scattered to form a scattered ray referred to as a secondary ray. The secondary ray formed by scattering travels in a path deviated from its original straight path and radiates on a surface of a pixel in the vicinity of the original pixel. Finally, besides the primary rays, the X-rays detected by each pixel of the detector further include the scattered secondary rays.
It can be known from the principle of formation of the secondary ray that a noise source may certainly be formed as a result of the secondary rays blending with the primary rays traveling in straight lines, causing reduction of the recognition capability of the detector on small difference of the contrast and decrease of the density resolution of the image system. Thus, it is a technical problem to be solved presently for those skilled in the art to reduce the scattered rays reaching the detector.
To solve the above technical problem, an X-ray collimator is often placed between the subject to be checked and the detector for reducing the influence of the scattered rays on the image density resolution.
Reference is made to FIG. 1, which is a schematic view of a conventional CT scanner having an X-ray collimator in the prior art.
A conventional X-ray collimator consists of a structure array which can absorb X-rays and is similar to the array of detector. A Chinese patent application No. 03826552.4 discloses an anti-scattered X-ray collimator of a CT scanning device. As shown in FIG. 1, the CT scanning device includes an X-ray source 2′, which may generate a conical X-ray beam 1′ indicated by dash lines in a controllable way, and an array 4′ of the X-ray detector 3′. The conical X-ray beam 1′ radiates from the focal spot 21′ of the X-ray source 2′. The array 4′ includes rows 31′ and columns 32′ of the X-ray detector 3′, wherein the rows 31′ are orthogonal or approximately orthogonal to the scanning center axis of the scanning device, and the columns 32′ are parallel or approximately parallel to the scanning center axis. The X-ray detector 3′ in the array 4′ is shielded by a two-dimensional X-ray collimator 5′. The X-ray collimator 5′ in FIG. 1 is partially cut away to illustrate the X-ray detector 3′. The X-ray collimator 5′ includes “row” slices 51′ positioned between the rows 31′ of the X-ray collimator 3′, and “column” slices 52′ positioned between the columns 32′ of the detector 3′. The “row” slices 51′ and “column” slices 52′ can be positioned such that their respective planes substantially intersect at the focal spot 21′.
With the above X-ray collimator 5′, the “raw” slices 51′ and “column” slices 52′ define and form a plurality of through holes or through gaps whose side walls are oriented in line with the straight paths from the focal spot 21′ to the surface of the X-ray detector 3′. That is to say, the extensions of the side walls of the plurality of through holes or through gaps intersect at the focal spot 21′. In this way, the area of the X-ray detector 3′ shielded by the side walls of the through holes or through gaps is smallest, i.e., the primary ray is shielded minimally, so that the primary rays may pass through the through holes or through gaps to the X-ray detector 3′ with least attenuation, thereby ensuring a high efficiency of the X-ray detector 3′.
As can be known from the above operating principle of the X-ray collimator 5′, there is a need for the X-ray collimator 5′ to absorb the scattered rays as many as possible on the one hand and to make most of the primary rays reach the surface of the X-ray detector 3′ with least attenuation on the other hand. The above requirement can be met by adjusting the height of the X-ray collimator 5′, the thickness of the walls of each through hole, and the shape of each through hole. And, it also has a high requirement for the size precision of each through hole in the X-ray collimator 5′. If the through holes cannot align with the pixel units of the X-ray detector 3′, on the one hand, the effective primary rays may be shielded, causing decrease of the detecting efficiency, and on the other hand, the non-uniform of the units may cause the image distortion. Therefore, the X-ray collimator 5′ plays an important role on a high-quality image scanned through a computed tomography system.
The current X-ray collimator as shown in FIG. 1 may be made by a first method in which sheet metal is stamped to form bended grid members and the grid members are adhered to each other, referring to the Chinese patent application No. 03826552.4 titled “Anti-scattered X-ray collimator of a CT scanning device” for further details. However, the collimator formed by the first method has a low precision due to springback during stamping. The X-ray collimator as shown in FIG. 1 can be made by a second method in which the slices with an array of grids are stacked, referring to the Chinese patent application No. 200810009502.5 titled “stacked CT collimator and the manufacturing method thereof” for further details. However, the stacked structure of multiple slices formed by the second method is complicate, and it is difficult to ensure uniformity of anti-scattering since the regions of slices where the primary rays pass through are required to have uniform sizes. The X-ray collimator can be made by a third method similar to 3D printing, referring to the Chinese patent applicant No. 02144468.4 titled “method of manufacturing scatting grid or collimator” for further details. However, the wall thickness of the through holes formed by printing in the third method may be thicker so as to greatly shield the primary rays, and may be non-uniform so as to cause non-uniformity of anti-scatting effect. The X-ray collimator may be formed by a fourth method of tenon connection, referring to the U.S. patent application with publication No. 20070258566A1 titled “ANTI-SCATTER-GRID” for further details. However, the X-ray collimator formed by the fourth method has a large number of components, causing the complicate assembling process.
Therefore, it is a technical problem to be solved presently for those skilled in the art to design an X-ray collimator of computed tomography system having an improved collimation effect and a simple structure.